1. Field of the Invention
The present invention relates to improved apparatus and methods for canceling feedback in audio systems such as hearing aids.
2. Description of the Prior Art
Mechanical and acoustic feedback limits the maximum gain that can be achieved in most hearing aids (Lybarger, S. F., "Acoustic feedback control", The Vanderbilt Hearing-Aid Report, Studebaker and Bess, Eds., Upper Darby, Pa.: Monographs in Contemporary Audiology, pp 87-90, 1982). System instability caused by feedback is sometimes audible as a continuous high-frequency tone or whistle emanating from the hearing aid. Mechanical vibrations from the receiver in a high-power hearing aid can be reduced by combining the outputs of two receivers mounted back-to-back so as to cancel the net mechanical moment; as much as 10 dB additional gain can be achieved before the onset of oscillation when this is done. But in most instruments, venting the BTE earmold or ITE shell establishes an acoustic feedback path that limits the maximum possible gain to less than 40 dB for a small vent and even less for large vents (Kates, J. M., "A computer simulation of hearing aid response and the effects of ear canal size", J. Acoust. Soc. Am., Vol. 83, pp 1952-1963, 1988). The acoustic feedback path includes the effects of the hearing-aid amplifier, receiver, and microphone as well as the vent acoustics.
The traditional procedure for increasing the stability of a hearing aid is to reduce the gain at high frequencies (Ammitzboll, K., "Resonant peak control", U.S. Pat. No. 4,689,818, 1987). Controlling feedback by modifying the system frequency response, however, means that the desired high-frequency response of the instrument must be sacrificed in order to maintain stability. Phase shifters and notch filters have also been tried (Egolf, D. P., "Review of the acoustic feedback literature from a control theory point of view", The Vanderbilt Hearing-Aid Report, Studebaker and Bess, Eds., Upper Darby, Pa.: Monographs in Contemporary Audiology, pp 94-103, 1982), but have not proven to be very effective.
A more effective technique is feedback cancellation, in which the feedback signal is estimated and subtracted from the microphone signal. Computer simulations and prototype digital systems indicate that increases in gain of between 6 and 17 dB can be achieved in an adaptive system before the onset of oscillation, and no loss of high-frequency response is observed (Bustamante, D. K., Worrell, T. L., and Williamson, M. J., "Measurement of adaptive suppression of acoustic feedback in hearing aids", Proc. 1989 Int. Conf. Acoust. Speech and Sig. Proc., Glasgow, pp 2017-2020, 1989; Engebretson, A. M., O'Connell, M. P., and Gong, F., "An adaptive feedback equalization algorithm for the CID digital hearing aid", Proc. 12th Annual Int. Conf. of the IEEE Eng. in Medicine and Biology Soc., Part 5, Philadelphia, Pa., pp 2286-2287, 1990; Kates, J. M., "Feedback cancellation in hearing aids: Results from a computer simulation", IEEE Trans. Sig. Proc., Vol.39, pp 553-562, 1991; Dyrlund, O., and Bisgaard, N., "Acoustic feedback margin improvements in hearing instruments using a prototype DFS (digital feedback suppression) system", Scand. Audiol., Vol. 20, pp 49-53, 1991; Engebretson, A. M., and French-St. George, M., "Properties of an adaptive feedback equalization algorithm", J. Rehab. Res. and Devel., Vol. 30, pp 8-16, 1993; Engebretson, A. M., O'Connell, M. P., and Zheng, B., "Electronic filters, hearing aids, and methods", U.S. Pat. No. 5,016,280; Williamson, M. J., and Bustamante, D. K., "Feedback suppression in digital signal processing hearing aids," U.S. Pat. No. 5,019,952).
In laboratory tests of a wearable digital hearing aid (French-St. George, M., Wood, D. J., and Engebretson, A. M., "Behavioral assessment of adaptive feedback cancellation in a digital hearing aid", J. Rehab. Res. and Devel., Vol. 30, pp 17-25, 1993), a group of hearing-impaired subjects used an additional 4 dB of gain when adaptive feedback cancellation was engaged and showed significantly better speech recognition in quiet and in a background of speech babble. Field trials of a feedback-cancellation system built into a BTE hearing aid have shown increases of 8-10 dB in the gain used by severely-impaired subjects (Bisgaard, N., "Digital feedback suppression: Clinical experiences with profoundly hearing impaired", In Recent Developments in Hearing Instrument Technology: 15th Danavox Symposium, Ed. by J. Beilin and G. R. Jensen, Kolding, Denmark, pp 370-384, 1993) and increases of 10-13 dB in the gain margin measured in real ears (Dyrlund, O., Henningsen, L. B., Bisgaard, N., and Jensen, J. H., "Digital feedback suppression (DFS): Characterization of feedback-margin improvements in a DFS hearing instrument", Scand. Audiol., Vol. 23, pp 135-138, 1994).
In some systems, the characteristics of the feedback path are estimated using a noise sequence continuously injected at a low level (Engebretson and French-St.George, 1993; Bisgaard, 1993, referenced above). The weight update of the adaptive filter also proceeds on a continuous basis, generally using the LMS algorithm (Widrow, B., McCool, J. M., Larimore, M. G., and Johnson, C. R., Jr., "Stationary and nonstationary learning characteristics of the LMS adaptive filter", Proc. IEEE, Vol. 64, pp 1151-1162, 1976). This approach results in a reduced SNR for the user due to the presence of the injected probe noise. In addition, the ability of the system to cancel the feedback may be reduced due to the presence of speech or ambient noise at the microphone input (Kates, 1991, referenced above; Maxwell, J. A., and Zurek, P. M., "Reducing acoustic feedback in hearing aids", IEEE Trans. Speech and Audio Proc., Vol. 3, pp 304-313, 1995). Better estimation of the feedback path will occur if the hearing-aid processing is turned off during the adaptation so that the instrument is operating in an open-loop rather than closed-loop mode while adaptation occurs (Kates, 1991). Furthermore, for a short noise burst used as the probe in an open-loop system, solving the Wiener-Hopf equation (Makhoul, J. "Linear prediction: A tutorial review," Proc. IEEE, Vol. 63, pp 561-580, 1975) for the optimum filter weights can result in greater feedback cancellation than found for LMS adaptation (Kates, 1991). For stationary conditions up to 7 dB of additional feedback cancellation is observed solving the Wiener-Hopf equation as compared to a continuously-adapting system, but this approach can have difficulty in tracking a changing acoustic environment because the weights are adapted only when a decision algorithm ascertains the need and the bursts of injected noise can be annoying (Maxwell and Zurek, 1995, referenced above).
A simpler approach is to use a fixed approximation to the feedback path instead of an adaptive filter. Levitt, H., Dugot, R. S., and Kopper, K. W., "Programmable digital hearing aid system", U.S. Pat. No. 4,731,850, 1988, proposed setting the feedback cancellation filter response when the hearing aid was fitted to the user. Woodruff, B. D., and Preves, D. A., "Fixed filter implementation of feedback cancellation for in-the-ear hearing aids", Proc. 1995 IEEE ASSP Workshop on Applications of Signal Processing to Audio and Acoustics, New Paltz, N.Y., paper 1.5, 1995, found that a feedback cancellation filter constructed from the average of the responses of 13 ears gave an improvement of 6-8 dB in maximum stable gain for an ITE instrument, while the optimum filter for each ear gave 9-11 dB improvement.
A need remains in the art for apparatus and methods to eliminate "whistling" due to feedback in unstable hearing-aids.